Nuclear medicine imaging is an extremely important tool used in the diagnosis and treatment of numerous different ailments. This generally involves injection of a radiopharmaceutical such as a technetium-99m complex or a iodine-131 complex which localizes in one or more areas or organs of the body. These radiopharmaceuticals will then emit gamma rays which can be detected to provide an image of the organ and/or tissue.
The Anger camera is the most commonly used image-formation device in nuclear medicine. The Anger camera uses a scintillation crystal, such as sodium iodide crystal, which absorbs gamma rays to produce light. An array of photomultiplier tubes adjacent to the crystal detects the light and amplifies the signal so as to indicate the location of the gamma-ray interaction and produce a two-dimensional plane-projected image of the object.
Anger cameras use lead collimators in order to eliminate all gamma rays excepting those which have a particular angle-of-incidence upon the camera face. The collimator, in effect, filters out the vast majority of gamma rays emitted from a source. But the collimator is required in order to achieve any usable spatial resolution.
Typically, with a state-of-the-art, triple-headed Anger camera, only one in 1,000 to 1,500 of emitted gamma rays are detected by the camera. The remainder are filtered out by the collimators. This requires an increased number of gamma rays in order to obtain a usable image. In order to achieve this, one can increase the radiation dose and/or increase the time for which the Anger camera is exposed to the source. Neither of these are particularly desirable.
Another problem, of course, with the lead collimators is weight. In certain applications particularly, very large plates of lead are required which may require special handling equipment such as booms and extra floor supports, drastically increasing the overall cost and decreasing the convenience of use.
A second method of imaging gamma rays in Nuclear Medicine is a Compton scatter camera. It has been suggested over the years to use Compton scattering to provide an image of a radiation source. Such a device is disclosed by Cree et al. (Cree, Michael J. and Philip J. Bones, Towards Direct Reconstruction from a Gamma Camera Based on Compton Scattering, IEEE TRANSACTIONS ON MEDICAL IMAGING, Vol. 13, No. 2, June 1994, pp. 398-407). Basically, a Compton scatter camera consists of two detector systems, herein called the primary and secondary (although some designs, e.g., multiple scatter cameras, do not make this distinction). The primary detector system, closest to the source, is designed so that the Compton scattering--that is, the scattering interaction between the gamma ray and an electron--is measured while the secondary detector system absorbs the scattered gamma rays and measures their energy. The idea is that a gamma ray emitted from the source is reasonably likely to undergo Compton scattering in the primary detector system, wherein the position and energy of the interaction are measured, then be absorbed in the secondary detector system wherein the position and energy of absorption are measured. From this, the location of the gamma-ray source can be determined, (after acquisition and reconstruction of data for numerous emissions).
FIG. 1 shows a diagrammatic depiction of these two types of interaction. For simplicity, the detector systems are depicted as planar arrays in FIG. 1. The angle of Compton scatter .theta. in the primary detector system can be calculated from the energy deposited in the primary detector .DELTA.E by ##EQU1## where E.sub.o is the initial gamma-ray energy, m is the electron mass, and c is the speed of light. This provides sufficient data to restrict the location of the original gamma-ray emission to a cone surface whose apex is centered at the location of the primary detector Compton scatter interaction, and whose axis is determined by the locations of the primary and secondary system interactions. Enough of these cone profiles will give sufficient information to reconstruct the original emission source distribution.
To be included in the reconstruction of the original radioisotope distribution, the Compton scatter camera minimally requires four pieces of information for each emitted gamma ray: (1) initial energy of the emitted gamma ray (E.sub.o); (2) location of the first Compton scatter interaction; (3) location of the second interaction (Compton scatter or photoelectric absorption); and (4) energy deposited by the first Compton scatter interaction (.DELTA.E). The following formula gives the resulting energy (E.sub.sc) of a gamma ray after it has Compton scattered off a free electron. ##EQU2##
FIG. 1 shows the geometry involved in a typical good event in a Compton scatter camera. One of the significant problems of an effective Compton camera is determining .DELTA.E. As indicated by Cree et al., the energy resolution in the primary (first) detector limits the resolution of scattering angle, and this angular resolution is itself dependent on the scattering angle. In the past, this has been resolved by providing primary detectors that have high energy resolution. Such high energy resolution detectors often require liquid nitrogen cooled semiconductor arrays which are edge-mounted to the thermal sink for minimal interference with transmitted and scattered gamma rays.
An array of detectors is required in order to optimally provide x, y and z position of the initial Compton scattering interaction in the primary detector system. This requires a very large physical surface area. Further, cryogenically cooling the primary detector system, which will be closest to the radiation source (the patient's body), provides significant difficulties. This must also be accomplished while, at the same time, providing a secondary detector system in sufficiently close physical proximity to the primary detector system to efficiently intercept and detect the Compton scattered gamma rays--that is, the absorption of the Compton scattered gamma rays by the secondary detector system.
Basically, this assumed requirement of high energy resolution in the primary detector system has posed technical difficulties which have prevented the commercialization of Compton scatter cameras for use in medical imaging. Although these devices have been suggested for medical imaging as, for example, in U.S. Pat. Nos. 5,175,434, 4,857,737 and 4,529,882 and the references cited therein, to date there is no commercial Compton scattering camera available for medical use. The Anger camera remains the primary workhorse for this application.